Efficient External Charger for an Implantable Medical Device Optimized for Fast Charging and Constrained by an Implant Power Dissipation Limit

ABSTRACT

An improved external charger for a battery in an implantable medical device (implant), and technique for charging the battery using such improved external charger, is disclosed. In one example, simulation data is used to model the power dissipation of the charging circuitry in the implant at varying levels of implant power. A power dissipation limit is chosen to constrain the charging circuitry from producing an inordinate amount of heat to the tissue surrounding the implant, and duty cycles are determined for the various levels of input intensities to ensure that the power limit is not exceeded. A maximum simulated average battery current determines the optimal (i.e., quickest) battery charging current, and at least an optimal value for a parameter indicative of that current, for example, the voltage across the battery charging circuitry, is determined and stored in the external charger. During charging, the actual value for that parameter is reported from the implant to the external charger, which in turn adjusts the intensity and/or duty cycle of the magnetic charging field consistent with the simulation to ensure that charging is as fast as possible, while still not exceeding the power dissipation limit.

FIELD OF THE INVENTION

The present invention relates generally to an external charger used toinductively charge a rechargeable battery within an implantable medicaldevice such as a neurostimulator.

BACKGROUND

Implantable stimulation devices generate and deliver electrical stimulito nerves and tissues for the therapy of various biological disorders,such as pacemakers to treat cardiac arrhythmia, defibrillators to treatcardiac fibrillation, cochlear stimulators to treat deafness, retinalstimulators to treat blindness, muscle stimulators to producecoordinated limb movement, spinal cord stimulators to treat chronicpain, cortical and deep brain stimulators to treat motor andpsychological disorders, occipital nerve stimulators to treat migraineheadaches, and other neural stimulators to treat urinary incontinence,sleep apnea, shoulder sublaxation, etc. The present invention may findapplicability in all such applications and in other implantable medicaldevice systems, although the description that follows will generallyfocus on the use of the invention in a Bion™ microstimulator devicesystem of the type disclosed in U.S. patent application Ser. No.12/425,505, filed Apr. 17, 2009.

Microstimulator devices typically comprise a small generally-cylindricalhousing which carries electrodes for producing a desired stimulationcurrent. Devices of this type are implanted proximate to the targettissue to allow the stimulation current to stimulate the target tissueto provide therapy for a wide variety of conditions and disorders. Amicrostimulator usually includes or carries stimulating electrodesintended to contact the patient's tissue, but may also have electrodescoupled to the body of the device via a lead or leads. A microstimulatormay have two or more electrodes. Microstimulators benefit fromsimplicity. Because of their small size, the microstimulator can bedirectly implanted at a site requiring patient therapy.

FIG. 1 illustrates an exemplary implantable microstimulator 100. Asshown, the microstimulator 100 includes a power source 145 such as abattery, a programmable memory 146, electrical circuitry 144, and a coil147. These components are housed within a capsule 202, which is usuallya thin, elongated cylinder, but may also be any other shape asdetermined by the structure of the desired target tissue, the method ofimplantation, the size and location of the power source 145 and/or thenumber and arrangement of external electrodes 142. In some embodiments,the volume of the capsule 202 is substantially equal to or less thanthree cubic centimeters.

The battery 145 supplies power to the various components within themicrostimulator 100, such the electrical circuitry 144 and the coil 147.The battery 145 also provides power for therapeutic stimulation currentsourced or sunk from the electrodes 142. The power source 145 may be aprimary battery, a rechargeable battery, a capacitor, or any othersuitable power source. Systems and methods for charging a rechargeablebattery 145 will be described further below.

The coil 147 is configured to receive and/or emit a magnetic field thatis used to communicate with, or receive power from, one or more externaldevices that support the implanted microstimulator 100, examples ofwhich will be described below. Such communication and/or power transfermay be transcutaneous as is well known.

The programmable memory 146 is used at least in part for storing one ormore sets of data, including electrical stimulation parameters that aresafe and efficacious for a particular medical condition and/or for aparticular patient. Electrical stimulation parameters control variousparameters of the stimulation current applied to a target tissueincluding, but not limited to, the frequency, pulse width, amplitude,burst pattern (e.g., burst on time and burst off time), duty cycle orburst repeat interval, ramp on time and ramp off time of the stimulationcurrent, etc.

The illustrated microstimulator 100 includes electrodes 142-1 and 142-2on the exterior of the capsule 202. The electrodes 142 may be disposedat either end of the capsule 202 as illustrated, or placed along thelength of the capsule. There may also be more than two electrodesarranged in an array along the length of the capsule. One of theelectrodes 142 may be designated as a stimulating electrode, with theother acting as an indifferent electrode (reference node) used tocomplete a stimulation circuit, producing monopolar stimulation. Or, oneelectrode may act as a cathode while the other acts as an anode,producing bipolar stimulation. Electrodes 142 may alternatively belocated at the ends of short, flexible leads. The use of such leadspermits, among other things, electrical stimulation to be directed totargeted tissue(s) a short distance from the surgical fixation of thebulk of the device 100.

The electrical circuitry 144 produces the electrical stimulation pulsesthat are delivered to the target nerve via the electrodes 142. Theelectrical circuitry 144 may include one or more microprocessors ormicrocontrollers configured to decode stimulation parameters from memory146 and generate the corresponding stimulation pulses. The electricalcircuitry 144 will generally also include other circuitry such as thecurrent source circuitry, the transmission and receiver circuitrycoupled to coil 147, electrode output capacitors, etc.

The external surfaces of the microstimulator 100 are preferably composedof biocompatible materials. For example, the capsule 202 may be made ofglass, ceramic, metal, or any other material that provides a hermeticpackage that excludes water but permits passage of the magnetic fieldsused to transmit data and/or power. The electrodes 142 may be made of anoble or refractory metal or compound, such as platinum, iridium,tantalum, titanium, titanium nitride, niobium or alloys of any of these,to avoid corrosion or electrolysis which could damage the surroundingtissues and the device.

The microstimulator 100 may also include one or more infusion outlets201, which facilitate the infusion of one or more drugs into the targettissue. Alternatively, catheters may be coupled to the infusion outlets201 to deliver the drug therapy to target tissue some distance from thebody of the microstimulator 100. If the microstimulator 100 isconfigured to provide a drug stimulation using infusion outlets 201, themicrostimulator 100 may also include a pump 149 that is configured tostore and dispense the one or more drugs.

Turning to FIG. 2, the microstimulator 100 is illustrated as implantedin a patient 150, and further shown are various external components thatmay be used to support the implanted microstimulator 100. An externalcontroller 155 may be used to program and test the microstimulator 100via communication link 156. Such link 156 is generally a two-way link,such that the microstimulator 100 can report its status or various otherparameters to the external controller 155. Communication on link 156occurs via magnetic inductive coupling. Thus, when data is to be sentfrom the external controller 155 to the microstimulator 100, a coil 158in the external controller 155 is excited to produce a magnetic fieldthat comprises the link 156, which magnetic field is detected at thecoil 147 in the microstimulator. Likewise, when data is to be sent fromthe microstimulator 100 to the external controller 155, the coil 147 isexcited to produce a magnetic field that comprises the link 156, whichmagnetic field is detected at the coil 158 in the external controller.Typically, the magnetic field is modulated, for example with FrequencyShift Keying (FSK) modulation or the like, to encode the data.

An external charger 151 provides power used to recharge the battery 145(FIG. 1). Such power transfer occurs by energizing the coil 157 in theexternal charger 151, which produces a magnetic field comprising link152. This magnetic field 152 energizes the coil 147 through the patient150's tissue, and which is rectified, filtered, and used to recharge thebattery 145 as explained further below. Link 152, like link 156, can bebidirectional to allow the microstimulator 100 to report statusinformation back to the external charger 151. For example, once thecircuitry 144 in the microstimulator 100 detects that the power source145 is fully charged, the coil 147 can signal that fact back to theexternal charger 151 so that charging can cease. Charging can occur atconvenient intervals for the patient 150, such as every night.

FIG. 3 illustrates salient portions of the microstimulator's powercircuitry 160. Charging energy (i.e., the magnetic charging field) isreceived at coil 147 via link 152. The coil 147 in combination withcapacitor 162 comprises a resonant circuit, or tank circuit, whichproduces an AC voltage at Va. This AC voltage is rectified by rectifiercircuitry 164, which can comprise a well-known 4-diode bridge circuit,although it is shown in FIG. 3 as a single diode for simplicity.Capacitor 166 assists to filter the signal at node Vb, such that Vb isessentially a DC voltage, although perhaps having a negligible ripple.Intervening between Vb and the rechargeable battery 145 is chargingcircuitry 170, which ultimately takes the DC voltage Vb and uses it toproduce a controlled battery charging current, Ibat. Charging circuitry170 is well known. One skilled in the art will recognize that the powercircuitry 160 may include other components not shown for simplicity.

It is generally desirable to charge the battery 145 as quickly aspossible to minimize inconvenience to the patient. One way to decreasecharging time is to increase the strength of the magnetic charging fieldby increasing the excitation current in the coil 157 of the externalcharger. Increasing the charging field will increase the current/voltageinduced in the coil 147 of the microstimulator 100, which increases thebattery charging current, Ibat. However, the strength of the magneticcharging field can only be increased so far before implant heatingbecomes a concern. One skilled in the art will understand that implantheating is an inevitable side effect of charging using magnetic fields.Heating can result from several different sources, such as eddy currentsin conductive portions of the implant, or heating of the variouscomponents in the power circuitry 160. Implant heating is a serioussafety concern; if an implant exceeds a given safe temperature (e.g.,41° C.), the tissue surrounding the implant may be aggravated ordamaged.

The art has recognized that heating can be controlled by controlling theintensity of the magnetic charging field produced at the externalcharger 151. For example, the current flowing through charging coil 157can be reduced to reduce the temperature of the implant during charging.The art has also recognized that heating can be regulated by dutycycling the charging field, i.e., by turning the charging field at theexternal charger 151 on and off. FIG. 4 generally shows the temperatureof the implant, T(IPG), for two different duty cycles, DC1 and DC2, fora given magnetic charging field. The first duty cycle, DC1, equals 50%,because the magnetic charging field is on for 50% of the time (i.e.,t1(on)=t1(off)). The second duty cycle, DC2, equals 75%, and hence themagnetic charger field stays on that much longer (i.e.,t2(on)=3t2(off)). As one would expect, higher duty cycles result inhigher temperatures in the implant: i.e., T1(IPG)<T2(IPG) as shown.

While changing the intensity or duty cycling of the magnetic chargingfield produced by the external charger 151 can be an effective means ofcontrolling implant temperature, the inventors have realized that suchapproaches do not adequately address important issues. First, knownprior approaches do not address whether the magnetic charging fieldintensity, duty cycle, or both, should be modified as a means oftemperature control. Moreover, such prior techniques are not understoodto consider efficient charging of the implant battery 145. Thus, one canchange the intensity and/or duty cycle of the magnetic charging field toarrive at suitable temperature control, but the particular parameterschosen may provide a charging power to the battery that is unnecessarilylow, which would prolong charging. Prolonged charging is inefficient,because that patient must wait an inordinate amount of time to fullycharge the battery 145 in his or her implant. Understandably, patientsdo not desire charging to take any longer than necessary.

Finding optimal charging conditions (intensity, duty cycle) thus remainsunknown with such prior art techniques, and this disclosure presents atechnique to combat this problem, and to make charging more efficientfrom both a time and implant heating perspective.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 illustrates a microstimulator implant, including a batteryrequiring periodical recharging from an external charger, in accordancewith the prior art.

FIG. 2 shows the implant in communication with, inter alia, an externalcharger in accordance with the prior art.

FIG. 3 illustrates charging circuitry within the implant in accordancewith the prior art.

FIG. 4 illustrates duty cycling the power at the external charger tocontrol implant temperature in accordance with the prior art.

FIGS. 5A-5C illustrate a simulation in accordance with one example ofthe disclosed technique.

FIG. 6 illustrates the relation between battery charging current (Ibat)and the voltage across the battery protection circuitry (Vnab) asrevealed from the simulation of FIG. 5.

FIG. 7 illustrates various duty cycles determined for the simulation ofFIG. 5 which will not exceed a prescribed power limit, and shows theapplication of such duty cycles on the power at the external chargingcoil (Iprim(rms)) and in the battery charging current (Ibat).

FIG. 8 illustrates the relation between the average battery chargingcurrent (Ibat(avg)) and Vnab as revealed from the simulation of FIG. 5,and shows the Vnab(opt) at which Ibat(avg) is maximized.

FIG. 9 illustrates storage of salient portions of the simulation toprepare the external charger for operation during an actual chargingsession, in accordance with one embodiment of the disclosed technique.

FIG. 10 illustrates circuitry in the external charger in accordance withone embodiment of the disclosed technique, including a memory storingsalient portions of the simulation relevant to optimization of thecharger's power parameters.

FIG. 11 illustrates a process in accordance with the disclosed techniquefor adjusting the power level and/or duty cycle of the power at theexternal charger in accordance with the stored simulation parameters.

FIG. 12 illustrates application of the disclosed technique assumingdefinition of an optimal range for Vnab(opt).

DETAILED DESCRIPTION

An improved external charger for a battery in an implantable medicaldevice (implant), and technique for charging the battery using suchimproved external charger, is disclosed. In one example, simulation datais used to model the power dissipation of the charging circuitry in theimplant at varying levels of implant power. A power dissipation limit ischosen to constrain the charging circuitry from producing an inordinateamount of heat to the tissue surrounding the implant, and duty cyclesare determined for the various levels of input intensities to ensurethat the power limit is not exceeded. A maximum simulated averagebattery current determines the optimal (i.e., quickest) battery chargingcurrent, and at least an optimal value for a parameter indicative ofthat current, for example, the voltage across the battery chargingcircuitry, is determined and stored in the external charger. Duringcharging, the actual value for that parameter is reported from theimplant to the external charger, which in turn adjusts the intensityand/or duty cycle of the magnetic charging field consistent with thesimulation to ensure that charging is as fast as possible, while stillnot exceeding the power dissipation limit. As a result, charging isoptimized to be as fast as possible, while still safe from a tissueheating perspective.

Prior to discussing the disclosed technique, reference is made to themicrostimulator power circuitry 160 of FIG. 3. While this circuitry isused in an explanation of the technique, it should be understood thatthe disclosed technique is not limited to use with the particular powercircuitry shown 160.

Various components in the power circuitry 160 within the implant willdraw power during the reception of a magnetic charging field from theexternal charger 151. In particular, the coil 147, its associated tankcapacitor 162, the rectification circuitry (diode) 164, chargingcircuitry 170, and the battery 145 itself will all dissipate power inthe form of heat. (Capacitor 166 will draw a comparatively negligibleamount of power, and thus is not further discussed). The sum total ofthe powers dissipated by each of these components must be consideredwhen understanding how the tissue surrounding the implant 100 will heatup during a charging session. For example, animal studies show that fora particular multiple-electrode microstimulator device, a radiated powerof 32 mW will raise the temperature of the tissue surrounding theimplant by approximately 4° C., while a total radiated power of 25.6 mWwill raise the temperature by 3.2° C. Of course, these values are onlyexemplary, and could vary; future values could be determined that aremore accurate, safer, etc. In any event, such animal studies havecorrelated power dissipation to tissue heating for a given implant.

An aspect of the disclosed technique seeks to keep the total dissipatedpower at or below a limit to ensure that the patient's tissue will notoverheat. Because a 4° C. rise in tissue temperature is generallyaccepted as safe for a patient, one example of the technique labors tokeep the total power dissipated from the power circuitry 160 at or below32 mW. Of course, different limits could be chosen, such as the 25.6mW/3.2° C. limit discussed above.

The inventors have noticed through simulations that power dissipationfrom the various components in the power circuitry 160 is complex andnon-linear in nature. One such simulation 200 is illustrated in FIGS.5A, 5B, and 5C. As will be discussed further below, certain portionsfrom simulation 200 are stored in the external charger 151 and will beused to regulate charging. However, before discussing a chargingoperation, simulation 200 is explained.

Simulation 200 shows the effect of varying the intensity (e.g., current)in the external controller's charging coil 157 (Iprim(rms)) on thevarious components in the power circuitry 160 of the implant 100, witheach successive row representing an increasing value for Iprim(rms).Because the simulation 200 results will vary depending on how full ordepleted the implant battery 145 is at a given moment, the depictedsimulation assumes a battery with a particular voltage of Vbat=3.1 V.Although not depicted, other simulations 200 at other battery voltages(e.g., 3.3V, 3.7V, 4.1V, etc.) may also be generated to provide accuratesimulation results as battery capacity starts to fill during charging.For example, if the battery 145 has a full capacity of Vbat=4.1V, thensimulations 200 may be generated for Vbat=3.1 V, 3.3V, 3.7V, and 4.1V tocover a range of expected battery capacity. However, if the variousparameters within simulation 200 do not vary appreciably with Vbat, thenthe generation of additional simulations 200 for different batterycapacities may not be necessary. A simulation program useful ingenerating a simulation 200 is Mentor Graphics Design Architect.

The simulation 200 assumes a particular coupling factor between theprimary coil 157 in the external charger 151 and the secondary coil 147in the implant 100, which coupling factor is modeled taking into accountfactors affecting such coupling, such as coil inductances, coilalignment, the distance and permittivity of any materials (e.g., tissue,air) between the coils, etc. In the depicted simulation, a couplingfactor k=0.017 was chosen to conservatively simulate a worst casealignment between the charging coils 157 and 147. In any event, thecoupling factor ultimately results in a simulated induced current incharging coil 147 in the implant (Isec(rms)), a current in theassociated tank capacitor 162 (Icap(rms)), a voltage across the coil 147(Vcoil(rms)), a DC voltage produced by the rectifier circuit (diode) 164(Vna), a battery charging current (Ibat), a battery voltage (Vbat)resulting from the input of the battery charging current, which batteryvoltage takes into account the internal resistance of the battery 145.Of course, relevant parameters for the various components in the powercircuitry 160 (resistances, capacitance, inductances, coupling factor,etc.) are input into the simulation program to allow it to generate thesimulation results.

Of particular interest in simulation 200 is the voltage across thecharging circuitry 170, Vnab, which represents the difference betweenVna and Vbat. Because the charging circuitry 170 is in line with thebattery charging current, Ibat, any voltage build up across the chargingcircuitry comprises undesired heat generation. Unfortunately, modelingshows that the amount of heat dissipation from the charging circuitry170 increases essentially exponentially as the battery charging currentincreases. This is shown in FIG. 6: as the battery charging current Ibatincreases, the voltage built up across the battery protection circuitryVnab increases at an increasingly fast rate. Because the powerdissipated by the charging circuit 170 equals the current times thevoltage, the power too essentially exponentially increases. In short,the parameter Vnab represents charging power wasted as heat, and as willbe seen below, is monitored and controlled in the disclosed technique topermit charging at an optimally efficient level.

From the various simulated voltages and currents in FIG. 5A, thesimulation 200 can further calculate the power dissipated by the variouscomponents in the power circuitry 160, as shown in FIG. 5B, which powersessentially comprise the product of the voltage across and currentthrough the various components. As shown, the power drawn by eachcomponent is represented by the element numeral for the component: forexample, the power drawn by the battery 145 during charging is denotedas P145. Pfes represents power drawn by front end switches in serieswith the charging circuitry 170, which switches are not depicted forsimplicity because their power dissipation are relatively small. The sumof the power dissipated by each of the components in the power circuitry160 is shown in the last column in FIG. 5B (Ptotal).

A review of the Ptotal parameter in simulation 200 illustrates a tissueheating concern for the designer. As discussed earlier, an acceptablelevel of total power dissipated by the power circuitry 160 should notexceed the 32 mW power dissipation limit in one example—a temperatureknown by experimentation to increase surrounding tissue by 4° C.However, all but the top three rows in FIG. 5B exceed this value (boldedfor easy viewing). In other words, simulation 200 shows that at higherexternal charger intensities (i.e., higher Iprim(rms)), the total heatgenerated in the implant 100 may be excessive.

One solution to keep the total power at or below 32 mW is to duty cyclethe power at the external charger 151, which computed duty cycle isshown in FIG. 5C. The duty cycle insures that the power dissipationlimit is not exceeded by dividing the limit (e.g., 32 mW) by thesimulated total power draw assuming no duty cycling (Ptotal).

The results of such duty cycling are shown in FIG. 7 for the third,fourth, and fifth rows in the simulation 200, i.e., when Iprim(rms)equals 600, 800, and 1000 mA. In the third row, the simulated totalpower dissipated was 27.5 mW, which is below the 32 mW limit. Hence,duty cycling would not be required for this level of input power (i.e.,for Iprim(rms)=600 mA). However, a duty cycle of 90% is imposed anywayto allow an off time, or telemetry window (TW), during which the implant100 can back-telemeter data to the external charger 151. The telemetrywindow (TW) may be 10 sec for example, meaning that the period for dutycycling is typically about 10 times larger, or 100 sec. While thetelemetry window TW can be fixed, it can also be made to vary dependingon how long is needed to send data back to the external charger 151. Forexample, the TW can be set to the exact time needed for datatransmission, with the on portion of the cycle similarly scaled to matchthe duty cycle required. A shorter duration for the total period of theduty cycle reduces ripple in the temperature of the implant 100.

As will be seen further below, it is advantageous to telemeter data(e.g., Vnab, Vbat) back to the external charger 151 to allow charging tobe iteratively optimized in real time. As can be seen in FIG. 7, thisduty cycle is imposed on the primary coil in the external charger(Iprim(rms)), which causes the same duty cycle in the battery chargingcurrent, Ibat. An average battery current, Ibat(avg), can be calculatedfrom the product of Ibat and the duty cycle to give an over-timeindication of the amount of charging current that is being received bythe battery, despite the duty cycling. The significance of Ibat(avg)will be discussed further below.

In the fourth row of the simulation 200 (Iprim(rms)=800 mA), thesimulated total power dissipated was 38.6 mW, above the 32 mW limit.Therefore, duty cycling is imposed as a heat control measure, inaddition to the desire for an off period to allow for back telemetry.Such duty cycling equals 82.9% (32/38.6) to ensure a total dissipatedpower of not more than 32 mW. The fifth row is similarly processed todetermine a duty cycle of 61.2%, and its effects on Iprim(rms) and Ibatare shown.

Additionally shown to the right in FIG. 5C are the computed duty cyclesfor the less-heat-intensive 25.6 mW/3.2° C. limit, which limit may bechosen to even further minimize patient discomfort or injury due to heatgeneration in the power circuitry 160. Again, the duty cycles arecomputed by dividing the limit (25.6 mW) by the simulated total powers(Ptotal).

Note from FIG. 7 that the average battery current, Ibat(avg), ismaximized when Iprim equals 800 mA. This average maximum,Ibat(avg)(opt)=11.6 mA, represents the optimal charging current for theimplant battery 145: it is the largest average current and hence willcharge the implant battery the fastest. Moreover, because of the dutycycling leading to the calculation of the Ibat(avg) values,Ibat(avg)(opt) is at the same time optimized to allow no more than 32 mWpower dissipation on average. Ibat(avg)(opt) is thus optimized for bothspeed and heat dissipation.

The disclosed technique seeks to maintain charging at this optimalaverage battery current. To so maintain Ibat(avg)(opt) during charging,it is useful to monitor a parameter indicative of the battery chargingcurrent, Ibat. One convenient parameter comprises Vnab, i.e., thevoltage that builds across the charging circuitry 170, although otherparameters indicative of the battery charging current could also be used(e.g., Vna). The Vnab parameter is easily measured in the implant, andas noted earlier represents wasted heat. FIG. 8 shows a graph ofIbat(avg) v. Vnab for the simulation 200 for the 32 mW/4° C. limit, andshows the maximum at 11.6 mA. The corresponding Vnab for this value,Vnab(opt) is 0.243 V (see fourth row, FIG. 5A). Vnab(opt) thusrepresents the voltage across the charging circuitry 170 that providesthe quickest charging of the implant battery 145, but which is safe froma heating perspective. As will be seen below, one implementation of thedisclosed technique is to maintain Vnab at Vnab(opt) during charging.

Prior to discussing use of the technique in an actual charging session,steps to this point in the process are summarized in FIG. 9, which stepslead to storing relevant parameters in the external charger 151. First,a power dissipation limit is chosen, such as the 32 mW/4° C. limitdiscussed previously. Then, the external charger 151/implant 100 systemis simulated to determine the relationship between Vnab and the dutycycle needed to stay compliant with the power dissipation limit. Thissimulation can occur assuming a particular battery voltage (Vbat) forthe battery 145 in the implant 100. Next, the relationship between Vnaband Ibat(avg) is determined using the duty cycle, and an optimalVnab(opt) is determined which corresponds to the maximum Ibat(avg).Thereafter, Vnab v. DC, and Vnab(opt) are stored in memory of theexternal charger, as will be discussed further shortly. Thereafter, thepreparation process repeats for a new battery voltage if necessary, butas noted earlier this may not be required if the various simulatedparameters do not vary strongly with Vbat.

FIG. 10 shows the external charger 151 as prepared with the parametersstored from FIG. 9. Shown with particularity is a memory 302, whichcontains at least a portion of the simulation 200, including the Vnab v.DC relationship for the 32 mW power dissipation limit and Vnab(opt) forVbat=3.1V. Also shown in part is the same information for Vbat=3.3V,although as just discussed this is not strictly necessary.Alternatively, memory 302 could contain the same information for otherpower dissipation limits (e.g., 25.6 mW/3.2° C.) as well, but this isnot shown for simplicity. The memory 302 containing these parameters iscoupled to (or could comprise part of) the microcontroller 300 in theexternal charger 151.

Also shown in FIG. 10 are the transmitter 304 and receiver 306 circuitscoupled to the external charger's coil 157, which circuitry is wellknown. The transmitter 304 produces an AC signal to cause the L-C tankcircuit (156/157) to resonate and in turn generate the magnetic chargingfield. As shown, the transmitter 304 receives control signals from themicrocontroller 300 to indicate the intensity (e.g., the magnitude ofIprim) and the duty cycle of the transmitter 304. The receiver 306receives data transmitted periodically from the implant 100, e.g.,during the telemetry window (TW) or off portions of the duty cycle (seeFIG. 7). Such data may be transmitted using radio-frequency (RF)telemetry, or Load Shift Keying (LSK) for example. (LSK is furtherdiscussed in U.S. patent application Ser. No. 12/354,406, filed Jan. 15,2009, for example).

Traditionally, such back telemetry from the implant to the externalcharger is used to transmit the capacity of the battery 145 duringcharging (Vbat), which informs the external charger 151 when the batteryis full and that charging can cease. Battery capacity is similarlyreported in disclosed system, but additionally, the Vnab value measuredat the implant 100 is also transmitted. Reporting of Vnab to theexternal charger 151 can take place at any suitable interval duringcharging, such as once every 100 seconds or so. The more frequently Vnabis reported, the more frequently charging can be optimized during thecharging session.

With the basic structure of the external charger 151 understood,attention can now focus on how charger 151 operates during an actualcharging session, which basic steps are shown in FIG. 11. First, theexternal charger 151 is turned on (e.g., by the patient), and generatesa magnetic charging field using an initial intensity level (i.e., aninitial Iprim) and an initial duty cycle. Simulation 200 does not helpmuch in determining initial values for the power and duty cycle levelsused at the external charger, as the coupling to the implant 100 duringa real charging session cannot be perfectly known in advance. Forexample, different patients may have their implants located at differentdepths in their tissues, or may have different physical alignmentsbetween their external chargers and their implants. In any event, theinitial power and duty cycle values are not important as they will bechanged in accordance with the disclosed technique as chargingprogresses, although they are logically set to initial values guaranteednot to injure the patient.

Periodically during charging, for example, perhaps every 100 seconds,the battery voltage (Vbat) and the voltage across the charging circuitry(Vnab) are measured at the implant 100, and telemetered to the externalcharger. Again, such telemetry can comprise RF or LSK telemetryperformed during the telemetry window (TW) or off periods in the dutycycle. How often to communicate, just like the time used forcommunication during the telemetry window (TW), may also be determinedby the length of the needed communication between implant and charger.Increasing the frequency of communication will reduce temperature ripplein the implant 100.

Once Vnab is reported, the microcontroller 300 consults memory 302 tosee if Vnab is optimal, i.e., if Vnab=Vnab(opt) for the reported Vbat.If not, intensity of the magnetic charging field is changed. Forexample, and referring to memory 300 in FIG. 10, if Vnab is near 0.293Vfor Vbat=3.1V, the microcontroller 300 would understand that theintensity is too high, and would reduce Iprim in an attempt to make Vnabapproach Vnab(opt). Conversely, if Vnab is near 0.181V, Iprim would beincreased.

At the same time, the duty cycle of the magnetic charging field wouldalso be changed to match the Vnab being reported. Modifying the dutycycle to match Vnab is important to ensure proper compliance with thepower dissipation limit. For example, and referring again to FIG. 10,assume again that Vnab is near 0.293V, but that the duty cycle currentlyimposed at the transmitter 304 is 85%. Reference to the storedparameters in memory 302 shows that this duty cycle is too high, andwill produce too much heat, i.e., more than the 32 mW power dissipationlimit. To keep the total dissipated power compliant with the limit, themicrocontroller 300, upon consulting memory 302, will change the dutycycle to 61.2%.

As shown in FIG. 11, once such intensity and duty cycle adjustments aremade at the external charger 151, the process repeats: Vbat and Vnab areagain reported after some time, and the intensity and duty cycleadjusted again if necessary. It should be noted that such iterativeadjustment of the power produced by the external charger 151 isparticularly helpful in applications where the coupling between theexternal charger 151 and the implant 100 might change. For example, thepatient may move the external charger relative to the implant during thecharging sessions. Such coupling changes can be compensated for usingthe disclosed technique, with adjustments made in situ to ensure thefastest charging within safe temperature limits.

To this point in the disclosure, it has been assumed that there is asingle optimal Vnab value, Vnab(opt). However, Vnab(opt) can alsorepresent a range of acceptable Vnab values. For example, the simulation200 in FIG. 5C shows three values for Ibat(avg) over 11 mA (rows fourthrough six), which correspond to Vnab values (FIG. 5A) of 0.243 to0.319V. Assuming that operation at any of these battery chargingcurrents provides satisfactorily quick charging of the implant battery145, Vnab(opt) can be set to a range between 0.243 to 0.319V, asillustrated in FIG. 12. Therefore, if Vnab is reported within thisrange, the intensity at the external charger (Iprim) would not bechanged. However, even if the intensity is not changed, it may still beprudent to vary the duty cycle in accordance with Vnab to ensurecompliance with the heat limit. In this regard, notice in FIG. 5C thatalthough Ibat(avg) does not change appreciably across the specified Vnabrange (from 11.6 to 11.0 mA), the duty cycle changes rather sharply(from 82.9 to 53.6%). However, depending on the particulars of thesimulation, and the conservative nature of the heat limit chosen,changing duty cycling within the Vnab(opt) range might not be necessary.In any event, defining Vnab(opt) as a range will simplify operation ofthe technique, and will require less frequent modification of themagnetic charging field at the external charger 151.

It should be understood that various parameters (e.g., Vnab(opt); a DCcorresponding to a particular Vnab) can be interpolated or extrapolatedfrom the simulation 200, and are therefore not necessarily constrainedto actual values appearing in the simulation. However, suchinterpolation was not shown to keep discussion of the technique simple.

Many of the parameters determined herein (e.g., Vnab(opt)) result fromthe simulation 200, which simulation provides a convenient expedient forunderstanding the external charger/implant system. However, not allimplementations of the technique will require the use of a simulation.Instead, empirical data, experimental models, direct analytical tools,or values chosen by other means, could be used depending uponconsideration of factors deemed important by the designer.

The disclosed technique limits the total power dissipated by theimplant. However, the technique can be constrained to control heating atonly a portion of the implant. For example, in larger implants orimplants with low heat conductivity, the technique can be employed tolimit the local heating at any section of the implant. In such anapplication, the technique can use a parameter (perhaps different fromVnab) indicative of heating to that section, and limiting heating ofthat particular section to tolerable limits. Thus, this modification tothe technique would only consider power dissipated as heat in therelevant section of the device.

Vnab is used in this disclosure as the measure indicative of excesspower dissipation. However, other parameters from the implant indicativeof incoming power and which can be used to control that power can alsobe used, such total power delivered to the battery, ripple of the coilvoltage, ripple of the rectified voltage, on time of the rectifyingcircuit, duty cycle of the rectifying circuit, etc. Of course, theseparameters could be measured or inferred in the implant in differentways.

Even though the technique describes the periodic measurement ofparameters in the implant during a charging session, and periodicadjustment of the magnetic charging field, “periodic” should not beunderstood as necessarily taking such actions at set intervals. Instead,“periodic” should be understood as taking a plurality of such actionsover time, even if not at set intervals.

While the inventions disclosed have been described by means of specificembodiments and applications thereof, numerous modifications andvariations could be made thereto by those skilled in the art withoutdeparting from the literal and equivalent scope of the inventions setforth in the claims.

1. A method for charging a battery in an implantable medical device,comprising: generating a magnetic charging field using an externalcharger, the magnetic charging field for producing a battery chargingcurrent in the implantable medical device during a charging session;periodically measuring during the charging session a parameter in theimplantable medical device; periodically telemetering during thecharging session the parameter to the external charger; and periodicallyadjusting during the charging session the magnetic charging field inaccordance with the parameter to maximize the produced battery chargingcurrent while not exceeding a maximum power dissipation limit for theimplantable medical device.
 2. The method of claim 1, wherein theparameter is indicative of the produced battery charging current.
 3. Themethod of claim 1, wherein adjusting the magnetic charging fieldcomprises adjusting its intensity or duty cycle.
 4. The method of claim1, wherein the parameter comprises a voltage across circuitry in linewith the battery charging current.
 5. The method of claim 4, wherein thecircuitry comprises charging circuitry serially coupled to the batteryin the implantable medical device.
 6. A method for charging a battery inan implantable medical device, comprising: generating a magneticcharging field using an external charger, the magnetic charging fieldfor producing a battery charging current in the implantable medicaldevice; measuring a parameter in the implantable medical deviceindicative of the produced battery charging current; telemetering theparameter to the external charger; comparing a value of the telemeteredparameter to an optimal value at the external charger, wherein theoptimal value indicates a maximum average battery current; and adjustingthe magnetic charging field in accordance with the comparison in amanner not to exceed a power dissipation limit for the implantablemedical device.
 7. The method of claim 6, wherein adjusting the magneticcharging field comprises adjusting its duty cycle, its intensity, orboth.
 8. The method of claim 6, wherein adjusting the magnetic chargingfield comprises adjusting an intensity of the magnetic charging field tobring the value of the telemetered parameter toward the optimal value.9. The method of claim 6, wherein the optimal value comprises a range ofvalues for the parameter.
 10. The method of claim 6, further comprisingtelemetering a voltage of the battery to the external charger, andwherein adjusting the magnetic charging field additionally occurs inaccordance with the telemetered battery voltage.
 11. A method forcharging a battery in an implantable medical device, comprising:generating a magnetic charging field using an external charger, themagnetic charging field for producing a battery charging current in theimplantable medical device; measuring a parameter in the implantablemedical device indicative of the produced battery charging current;telemetering the parameter to the external charger; consulting datastored in the external charger which relates values of the parameter toa duty cycle for the external charger; determining a duty cycle from thedata for a value of the telemetered parameter; and duty cycling themagnetic charging field in accordance with the determined duty cycle.12. The method of claim 11, wherein the duty cycle is determined toensure that the implantable medical device does not exceed a powerdissipation limit.
 13. The method of claim 11, wherein the parametercomprise a voltage across charging circuitry in line with the battery.14. The method of claim 11, wherein an intensity of the magneticcharging field is also varied in accordance with the telemeteredparameter.
 15. The method of claim 11, wherein the parameter istelemetered to the external charger by Load Shift Keying in theimplantable medical device.
 16. A method for charging a battery in animplantable medical device, comprising: generating a magnetic chargingfield using an external charger, the magnetic charging field forproducing a battery charging current in the implantable medical device;measuring a parameter in the implantable medical device; telemeteringthe parameter to the external charger; and adjusting an intensity andduty cycle of the magnetic charging field in response to the telemeteredparameter, wherein the adjustment ensures a power dissipation limit forthe implantable medical device is not exceeded.
 17. The method of claim16, wherein the parameter is telemetered to the external charger duringan off cycle in the duty cycle of the magnetic charging field.
 18. Themethod of claim 16, wherein the method repeats itself to iterativelyadjust the intensity and duty cycle of the magnetic charging field. 19.The method of claim 16, wherein the parameter is indicative of heatingin the implantable medical device during charging.
 20. (canceled) 21.The method of claim 20 16, wherein the adjustment additionally ensuresthat the produced battery charging current is maximized for the powerdissipation limit.
 22. An external charger designed to ensure compliancewith a power dissipation limit for an implantable medical device duringcharging of a battery in the implantable medical device, comprising: amemory in which is stored: a first relationship between values for (i) aparameter measureable in the implant, and (ii) a duty cycle for theexternal charger, wherein the duty cycles ensure that the powerdissipation limit is not exceeded; and a first optimal value for theparameter in the external charger, wherein the first optimal value isindicative of a maximum battery charging current.
 23. The externalcharger of claim 22, further comprising a transmitter for broadcasting amagnetic charging field to the implantable medical device, and areceiver for receiving the parameter from the implantable medicaldevice.
 24. The external charger of claim 23, further comprising amicrocontroller in communication with the memory, the transmitter, andthe receiver.
 25. The external charger of claim 22, wherein thetransmitter implements controls of a duty cycle and intensity of themagnetic charging field.
 26. The external charger of claim 22, furthercomprising another data structure for a second battery voltage differentfrom the first battery voltage.
 27. The external charger of claim 22,wherein the memory further stores: a second relationship between valuesfor (i) the parameter measureable in the implant, and (ii) a duty cyclefor the external charger, wherein the duty cycles ensure that the powerdissipation limit is not exceeded; and a second optimal value for theparameter in the external charger, wherein the second optimal value isindicative of a maximum battery charging current.
 28. The externalcharger of claim 27, wherein the first relationship and first optimalvalue correspond to a first battery voltage, and wherein the secondrelationship and second optimal value correspond to a second batteryvoltage.
 29. A method for programming an external charger to ensurecompliance with a power dissipation limit for an implantable medicaldevice during charging, comprising: determining an external charger dutycycle and an average battery charging current for a plurality ofexternal charger intensities, wherein the duty cycles are determinedsuch that the implantable medical device does not exceed the powerdissipation limit; determining a parameter measureable in theimplantable medical device which is indicative of the average batterycharging current; determining a maximum average battery current and acorresponding optimal value for the parameter; storing a relationshipbetween the parameter and the duty cycles in the external charger; andstoring the optimal value for the parameter in the external charger. 30.The method of claim 29, wherein the determined values are determined bya computer simulation.
 31. The method of claim 29, wherein the parameteris indicative of the battery charging current.
 32. The method of claim29, wherein the parameter is indicative of heat generation in theimplantable medical device.